Method of improving electrode tissue interface

ABSTRACT

A critical element of a retinal prosthesis is the stimulating electrode array, which is placed in close proximity to the retina. It is via this interface that a retinal prosthesis electrically stimulates nerve cells to produce the perception of light. The impedance load seen by the current driver consists of the tissue resistance and the complex electrode impedance. The results show that the tissue resistance of the retina is significantly greater than that of the vitreous humor in the eye. Circuit models of the electrode-retina interface are used to parameterize the different contributors to the overall impedance.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application claims benefit of U.S. Provisional Patent applicationSer. No. 60/860,388, filed on Nov. 20, 2006, entitled “ElectricalProperties of Retinal-Electrode Interface,” the disclosure of which isincorporated herein by reference.

GOVERNMENT RIGHTS NOTICE

This invention was made with government support under grant No.R24EY12893-01. The government has certain rights in the invention.

FIELD OF THE INVENTION

The present invention relates to a method of improving the electrodetissue interface in a visual prosthesis.

BACKGROUND OF THE INVENTION

In 1755 LeRoy passed the discharge of a Leyden jar through the orbit ofa man who was blind from cataract and the patient saw “flames passingrapidly downwards.” Ever since, there has been a fascination withelectrically elicited visual perception. The general concept ofelectrical stimulation of retinal cells to produce these flashes oflight or phosphenes has been known for quite some time. Based on thesegeneral principles, some early attempts at devising prosthesis foraiding the visually impaired have included attaching electrodes to thehead or eyelids of patients. While some of these early attempts met withsome limited success, these early prosthetic devices were large, bulkyand could not produce adequate simulated vision to truly aid thevisually impaired.

In the early 1930's, Foerster investigated the effect of electricallystimulating the exposed occipital pole of one cerebral hemisphere. Hefound that, when a point at the extreme occipital pole was stimulated,the patient perceived a small spot of light directly in front andmotionless (a phosphene). Subsequently, Brindley and Lewin (1968)thoroughly studied electrical stimulation of the human occipital(visual) cortex. By varying the stimulation parameters, theseinvestigators described in detail the location of the phosphenesproduced relative to the specific region of the occipital cortexstimulated. These experiments demonstrated: (1) the consistent shape andposition of phosphenes; (2) that increased stimulation pulse durationmade phosphenes brighter; and (3) that there was no detectableinteraction between neighboring electrodes which were as close as 1 mm-5mm, preferably 2 mm-3 mm, more preferably about 2.4 mm apart.

As intraocular surgical techniques have advanced, it has become possibleto apply stimulation on small groups and even on individual retinalcells to generate focused phosphenes through devices implanted withinthe eye itself. This has sparked renewed interest in developing methodsand apparatuses to aid the visually impaired. Specifically, great efforthas been expended in the area of intraocular visual prosthesis devicesin an effort to restore vision in cases where blindness is caused byphotoreceptor degenerative retinal diseases such as retinitis pigmentosaand age related macular degeneration which affect millions of peopleworldwide.

Neural tissue can be artificially stimulated and activated by prostheticdevices that pass pulses of electrical current through electrodes onsuch a device. The passage of current causes changes in electricalpotentials across visual neuronal membranes, which can initiate visualneuron action potentials, which are the means of information transfer inthe nervous system.

Based on this mechanism, it is possible to input information into thenervous system by coding the information as a sequence of electricalpulses which are relayed to the nervous system via the prostheticdevice. In this way, it is possible to provide artificial sensationsincluding vision.

One typical application of neural tissue stimulation is in therehabilitation of the blind. Some forms of blindness involve selectiveloss of the light sensitive transducers of the retina. Other retinalneurons remain viable, however, and may be activated in the mannerdescribed above by placement of a prosthetic electrode device on theinner (toward the vitreous) retinal surface (epiretial). This placementmust be mechanically stable, minimize the distance between the deviceelectrodes and the visual neurons, and avoid undue compression of thevisual neurons.

In 1986, Bullara (U.S. Pat. No. 4,573,481) patented an electrodeassembly for surgical implantation on a nerve. The matrix was siliconewith embedded iridium electrodes. The assembly fit around a nerve tostimulate it.

Dawson and Radtke stimulated cat's retina by direct electricalstimulation of the retinal ganglion cell layer. These experimentersplaced nine and then fourteen electrodes upon the inner retinal layer(i.e., primarily the ganglion cell layer) of two cats. Their experimentssuggested that electrical stimulation of the retina with 30 to 100 uAcurrent resulted in visual cortical responses. These experiments werecarried out with needle-shaped electrodes that penetrated the surface ofthe retina (see also U.S. Pat. No. 4,628,933 to Michelson).

The Michelson '933 apparatus includes an array of photosensitive deviceson its surface that are connected to a plurality of electrodespositioned on the opposite surface of the device to stimulate theretina. These electrodes are disposed to form an array similar to a “bedof nails” having conductors which impinge directly on the retina tostimulate the retinal cells. U.S. Pat. No. 4,837,049 to Byers describesspike electrodes for neural stimulation. Each spike electrode piercesneural tissue for better electrical contact. U.S. Pat. No. 5,215,088 toNorman describes an array of spike electrodes for cortical stimulation.Each spike pierces cortical tissue for better electrical contact.

The art of implanting an intraocular prosthetic device to electricallystimulate the retina was advanced with the introduction of retinal tacksin retinal surgery. De Juan, et al. at Duke University Eye Centerinserted retinal tacks into retinas in an effort to reattach retinasthat had detached from the underlying choroid, which is the source ofblood supply for the outer retina and thus the photoreceptors. See,e.g., E. de Juan, et al., 99 Am. J. Opthalmol. 272 (1985). These retinaltacks have proved to be biocompatible and remain embedded in the retina,and choroid/sclera, effectively pinning the retina against the choroidand the posterior aspects of the globe. Retinal tacks are one way toattach a retinal array to the retina. U.S. Pat. No. 5,109,844 to de Juandescribes a flat electrode array placed against the retina for visualstimulation. U.S. Pat. No. 5,935,155 to Humayun describes a visualprosthesis for use with the flat retinal array described in de Juan.

SUMMARY OF THE INVENTION

A critical element of a retinal prosthesis is the stimulating electrodearray, which is placed in close proximity to the retina. It is via thisinterface that a retinal prosthesis electrically stimulates nerve cellsto produce the perception of light. The impedance load seen by thecurrent driver consists of the tissue resistance and the complexelectrode impedance. The results show that the tissue resistance of theretina is significantly greater than that of the vitreous humor in theeye. Circuit models of the electrode-retina interface are used toparameterize the different contributors to the overall impedance. Theinvention involves a method of optimizing an electrode tissue interfacewhich comprises placing an electrode in the vicinity of neural tissue;stimulating the neural tissue through said electrode; measuring theimpedance of the electrode tissue interface; and altering the locationof said electrode based on said impedance measurement.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a diagram of the preferred electrode array layout.

FIG. 2 is a schematic representation of the tissue electrode interface.

FIG. 3-1 is a set of graphs showing the relationship of impedancemodulus, phase and electrode diameter.

FIG. 3-2 is a set of graphs showing the relationship of impedancemodulus, phase and electrode diameter.

FIG. 4 is a set of graphs showing the relationship of impedance modulusvs. electrode diameter in vitreous and on the retina.

FIG. 5 is a bar graph showing the relationship of impedance modulus vs.frequency at various electrode diameters.

FIG. 6 shows an average impedance measurement of a patient.

FIG. 7 shows RS and polarization impedance electrodes against theretina.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

Retinal prostheses are implantable electronic devices designed toelectrically stimulate retinal neurons when the photoreceptors of theretina are absent, due to photoreceptor degenerative disease.Age-related macular degeneration and retinitis pigmentosa are two suchdiseases that blind millions worldwide. Several clinical trials ofprototype retinal prostheses are currently on-going. All have shown theability to evoke phosphenes (spots of light) in subjects otherwiseblind. The systems are similar in their implementation. They use anexternal camera to capture light information, an external processor tocode the image information and then wirelessly transmit stimuluscommands to implants. The implants have custom circuits which decode thestimulus command signal and output electrical charge via a controlledcurrent or voltage signal. The prototype devices all use electricalstimulation of retinal cells with extracellular electrodes as the meansof creating a sensation of light in blind individuals.

The stimulating electrode is an important consideration when designing aretinal prosthesis. The impedance of the electrode and retina largelydefine the output load for the stimulator circuit, and thus the powerconsumption of the implanted system. It is desirable to have a denseinterface, i.e. to put as many electrodes as possible in a small area ofthe retina, to achieve maximum visual acuity. This suggests that smalldiameter electrodes are preferable, but smaller electrodes have a numberof negative consequences. These include higher impedance, which, for agiven amount of stimulating current, requires a larger power supply forthe stimulator chip and high charge density. Even if a charge balancedstimulus is used to make ensure no DC current, it is still possible todamage tissue if excessive charge density is used to stimulate. To guidethe design of a retinal prosthesis, it is important to know theelectrical interface properties. In this disclosure, the intraocularimpedance is studied as a function of electrode location in the eye andelectrode diameter.

The eye model used for these experiments was a cadaveric porcine eye,with an artificial cornea attached to allow viewing for retinal surgery.A smaller data set with in situ canine eye has been known, however ithas been switched to cadaveric porcine eye for several reasons,including: 1) comparable results to the live canine eye 2) ease of useduring implantation and surgery 3) increased availability, since theopportunity to perform canine experiments came only during terminalsurgeries that involved other experiments. Several electrode arraydesigns were tested. Features common to all designs include polyimideinsulating layers as disclosed in U.S. patent applications Nos.20060247754, 20060259112, and 20070265665, and circular, platinumelectrodes as disclosed in U.S. Pat. Nos. 6,974,533 and 7,181,287 andU.S. patent applications Nos. 20020111658, 20070173905, 20070026048,20030195601, 20060259108, 20050271895, 20060063062, 20040220652,20060259109, 20070092786, 20070089994, 20070089992, 20070092750, and20070191911, the disclosure of which are incorporated herein byreference.

The electrodes and insulating layers were coplanar. The electrodediameters tested include 50, 100, and 200 μm. Two array designs wereused. Both arrays had a 4×4 arrangement of electrodes. One array hadonly 200 μm electrodes. The other array had 10, 25, 50, and 100 μmelectrodes. The 10 and 25 μm diameter electrodes did not yieldconsistent results so these sizes were not included in further testing.The layout of the multi-size electrode is shown in FIG. 1. FIG. 1 showselectrode array layout used for testing. The electrode spacing was 800μm center to center in the horizontal and vertical direction. Anelectrode with a similar layout, but with 200 μm diameter electrodes wasalso used.

For the cadaveric porcine eye (eye) surgery the eye was fixed in aStyrofoam holder. An infusion line was placed to maintain the pressureof the eye. Since the cornea and lens opacify shortly after enucleation,a false cornea (keratoprosthesis) was used to enable clear viewing ofthe vitreous cavity. A trephine was used to open a circular window inthe cornea and the crystalline lens was removed through this opening.The keratoprosthesis was sutured into place in the window in the cornea.A standard three-port vitrectomy was performed to replace the vitreousgel with balanced salt solution (BSS), which is a standard vitreousreplacement solution. To ensure that the electrode was in close contactwith the retina, care was taken to remove as much vitreous as possiblefrom the retinal surface. One sclerotomy wound was increased to 5 mm toaccommodate insertion of the electrode array. After array insertion, thesclerotomies were closed so that minimal fluid was leaking from the eye,as determined by the drip rate of the infusion line. The electrode arraywas placed in the middle of the vitreous cavity for the first set ofmeasurements. Then the electrode was placed on the retina for the secondset of measurements. When placing the array on the retina, a retinaltack was used to ensure close placement.

AC impedance was measured using a commercial potentiostat (FAS1, GamryInstruments, Inc.). The platinum test electrode was connected to theworking electrode input and a large gold disc electrode (1 cm diameter,two sided) served as the counter and return electrode. The counterelectrode was placed in contact with outside of the eye, on the oppositeside of the eye from the large sclerotomy. Impedance was measured at 5points/decade from 10 Hz to 100 kHz. The excitation signal was a 10 mVsine wave. No DC bias potential was applied so the electrode was at opencircuit potential.

Circuit models were used to parameterize the different parts of theelectrode-tissue impedance. A common model used to describe theelectrode-electrolyte impedance spectrum is the three element modelshown in FIG. 2. FIG. 2 shows a three element circuit model. R_(u)represents tissue resistance, C_(F) represents double layer capacitance,and R_(p) represents Faradaic charge transfer. For this modeling, a CPEwas substitute for C_(F) and a parallel capacitance was added to accountfor parasitic system capacitance. Charge injection via the electrode cantake place either via charging and discharging across the double layercapacitive layer represented by capacitance C of the model or throughelectrochemical reactions across the interface modeled by a parallelresistor. The electrolyte is modeled by a series resistance. Variationson this include use of a constant phase element (CPE) in place of theparallel capacitor when the electrode is roughened, as was the case inthese studies. The CPE represents a network of resistors and capacitors,similar to a transmission line, and its impedance is expressed asZ_(CPE)=Y_(o)(jω)^(−n)). If n is close to 1, then the CPE is primarilycapacitive, if n is close to 0 then the CPE is primarily resistive.Another element was included in the model to account for experimentalconditions. A parasitic capacitance, due to long lead wires, wasincluded in parallel with the entire model. This was evident ascapacitive impedance at high frequency. Models were constructed and fitusing software from Gamry, Inc.

FIG. 3 depict the impedance and phase as a function of frequency for thedifferent sized electrodes when the device was in the vitreous cavityversus tacked to retina. In general, the plots are similar to othermicroelectrode impedance data. FIG. 3 shows impedance modulus (leftcolumn) and phase (right column) for 50 (top), 100 (middle) and 200(bottom) μm diameter electrodes in the vitreous and on the retina. Theimpedance modulus is higher when the device is on the retina versus whenthe device is in the vitreous, but only at higher frequency. In thisfrequency region, the impedance modulus is not highly dependent onfrequency, suggesting more resistive impedance. With decreasingfrequency, impedance modulus becomes more frequency dependent,suggesting capacitive impedance in this region. Similarly, the phase isdifferent at higher frequencies but converges to a capacitive phase atlower frequencies. Modeling results suggest that the capacitive phase atthe highest frequency, when the device was on the retina for 50 μm and100 μm electrodes, was due to parasitic capacitance inherent in themeasurement. Impedance increased with decreasing electrode size for bothvitreous and retina electrodes (FIG. 4). FIG. 4 shows especiallyimpedance modulus versus size in vitreous (left) and on retina.

FIG. 5 shows the ratio between impedance modulus when the electrode isagainst the retina versus impedance modulus when the electrode array isin mid vitreous. The average impedance modulus values for all theelectrodes of a particular size were used for the ratios. The ratiosindicate that the impedance modulus is between 3 and 6 times higher fromthe 50 and 100 μm diameter electrodes when the electrode is on thesurface of the retina. The 200 μm diameter electrode impedance modulusincreased less dramatically but still increased between 1.5 and 2 times.These increases were noted at 100 kHz, 10 kHz, and 1 kHz, but by 100 Hz,the impedance modulus was virtually unchanged between vitreous andretina electrodes. FIG. 5 especially shows—Ratio of Zmod [retina]/Zmod[vitreous] at 4 frequencies.

Since electrical stimulation uses a high frequency signal, the impedancedata was limited to 10 Hz-100 kHz. In this range, the electrode-tissuecircuit model reduces to two elements in series: the tissue resistanceand the CPE. To account for parasitic capacitance in the long lead wire,a parallel capacitor was added across the entire model. The averageparameter values for each model element are listed in table I. Theaverage series resistance (R_(u)) was 8.6, 6.7, and 2 times larger onthe retina for the 50, 100, and 200 μm electrodes, respectively. The CPEmagnitude Y_(o) was not as position dependent and the CPE exponent n isclose to 1, indicating a mostly capacitive CPE. The parasiticcapacitance was independent of the test conditions, suggesting asystemic capacitance.

TABLE 1 Circuit model parameters predicted from Impedance and Pulse DataY₀ R_(u) (Kilo Ohms) (nano C_(parasitic) Size (Location) from EISSiemens) n (pico Farads) 50 (retina) 74.35 29.975 0.824 99.0125 100(retina) 30.69 99.575 0.842 105.546 200 (retina) 5.415 388.623 0.887117.2 50 (vitreous) 8.623 24.245 0.856 150.075 100 (vitreous) 4.59666.511 0.907 138.91 200 (vitreous) 2.749 319.15 0.905 125.59

FIG. 6 is an impedance component of A60 Patient data—Rs and Zp. Theincrease of overall impedance due to tissue is mostly due to increase ofthe Rs component of impedance.

FIG. 7 is an overall impedance increase of electrodes against retinaltissue shows a great contribution from series resistance componentcompared to impedance from polarization impedance.

Impedance measurement is a method of monitoring the electrical status ofan electrode and its adjacent electrolyte environments. Afterimplantation, some electrodes show an increase in overall impedance. Byanalyzing the voltage waveform data, one can analyze the overallimpedance in terms of its component impedances: series resistance andpolarization impedance. Series resistance (Rs) is a measure of theadjacent tissue environment, while the polarization impedance is ameasure of the surface quality of the electrode. The current data fromsix A60 patients suggest that the increase in impedance afterimplantation is mainly due to an increase in the series resistancecomponent of the overall impedance and less from the polarizationimpedance.

OCT data from six A60 patients show 20 electrodes that are in directcontact with retinal tissue. The voltage waveform data of these 20electrodes were analyzed to determine the series resistance (Rs) andpolarization impedance (Zp) components of the overall impedance (Zt).Comparison with series resistance and polarization impedance ofelectrodes not in contact with retina shows that the overall impedanceincrease of the electrodes against retina have a greater component fromthe increase in series resistance, see FIG. 6. Similarly, when data isanalyzed by individual electrodes in contact with the retina tissue,increase in overall impedance is seen to be directly correlated withincrease in Rs, see FIG. 7.

Neural stimulator design must consider the tissue load impedance as partof the design because of Ohm's law V=IR. For a given V set by the chipvoltage supply, the tissue impedance R will limit the current output I.While designing a chip with a large voltage supply is one strategy toensure sufficient I, the cost of that approach is higher powerconsumption, a larger chip, and possible thermal dissipation issues inthe body. Usually it is noted a limit of 20 mW thermal dissipation for aresistive load placed on the retina. There have been modeled chipsoperating in the eye and have noted similar power limits. It is alsopossible to limit I, by increasing the pulse duration of the stimulus.There may be perceptual advantages to longer stimulation and lowercompliance voltages can also be used. However, longer pulses areinefficient for neural stimulation compared to shorter pulses, so theenergy consumption of the device would increase, requiring larger powertelemetry components. Thus, a thorough characterization of the loadimpedance for the neural stimulator is needed to optimize the interfaceand the overall system.

It was noted that the resistance of the retina was greater than that ofthe vitreous. This is in agreement with other studies that correlatedpsychophysical threshold and electrode impedance in humans withprototype retinal implants. Those studies found a trend towards higherimpedance for electrodes with lower thresholds. Both higher impedanceand lower thresholds suggest that an electrode is closer to the retina.Thus, there are positive and negative consequences of having theelectrode near the retina. A higher resistance will increase powerconsumption but a lower threshold requirement will decrease powerconsumption. Since it is also desirable to limit the area of retinaactivated for higher acuity, a closer proximity for the electrode isbest and stimulation systems will need to take into account higherimpedance. The reasons for higher retinal impedance versus vitreous canbe found in the anatomy. In general, neural tissue has higherresistivity than saline. However, the retina has unique structures thatmay further increase the impedance. The retina, if considered as amaterial, can be described as a laminated structure with each layerhaving different electrical conductivity. In particular, the innerlimiting membrane and the retina pigment epithelium have been found tohave higher resistivity than other parts of the retina.

Circuit models can provide useful information for system design ofneural stimulators. With respect to the tissue resistance, the dataobtained from small signal impedance data, obtained using a 10 mV sinewave, can be reasonably extrapolated to the large signal response, wherethe voltage across the tissue can be several Volts. However, theelectrode capacitance in general increases with increasing currentdensity, so care must be taken when using circuit models from impedancedata as model loads for circuit design. Impedance data willunderestimate the available capacitance for stimulation. Therefore,using impedance data to estimate the required voltage for a stimulatorchip may result in higher voltage than is actually needed.

Electrode impedance is strongly influenced by the location of theelectrode in the eye. An electrode in close apposition to the retinawill have significantly higher impedance. This will increase the powerrequirements for an implanted stimulator. Understanding the electricalinterface between a retinal prosthesis and the retina is an importantpart of optimizing the interface and the implanted system.

Accordingly, what has been shown is an improved method making a hermeticpackage for implantation in a body. While the invention has beendescribed by means of specific embodiments and applications thereof, itis understood that numerous modifications and variations could be madethereto by those skilled in the art without departing from the spiritand scope of the invention. It is therefore to be understood that withinthe scope of the claims, the invention may be practiced otherwise thanas specifically described herein.

1. A method of optimizing an electrode tissue interface comprising:placing at least one electrode in the vicinity of neural tissue;stimulating the neural tissue through the electrode; measuring theimpedance of the electrode tissue interface; and altering the locationof said electrode based on said impedance measurement.
 2. The method ofclaim 1 wherein the distance between neighboring electrodes is from 1mm-5 mm.
 3. The method of claim 2 wherein the distance betweenneighboring electrodes is from 2 mm-3 mm.
 4. The method of claim 3wherein the distance between neighboring electrodes is from 2.2 mm-2.5mm.
 5. The method of claim 1 wherein the electrode diameters are from 10μm to 250 μm.
 6. The method of claim 5 wherein the electrode diametersare from 40 μm to 120 μm.
 7. The method of claim 1 wherein the electrodespacing is from 600 μm to 1000 μm center to center in the horizontal andvertical direction.
 8. The method of claim 1 wherein the electrodespacing is from 850 μm to 950 μm center to center in the horizontal andvertical direction.
 9. The method of claim 1 wherein sclerotomy woundwas increased for 3 mm to 7 mm to accommodate insertion of the electrodearray.
 10. The method of claim 1 wherein the impedance data is limitedto a range from 10 Hz to 100 kHz.
 11. The method of claim 1 wherein theseries resistance (Ru) is in the range from 1 kilo ohm to 80 kilo ohms.12. The method of claim 1 wherein the CPE magnitude is in the range from20 nano siemens to 400 nano siemens.